1. Field of the Invention
The invention discloses a tubing for pumps, said tubing having at least one longitudinal relatively thin wall, that provides means to measure inlet and outlet pressure of pumped fluid, without contacting the fluid, and to regulate flow as a function of inlet outlet pressure of the pump. A peristaltic pump designed to take advantage of the special properties of the thin wall tubing is also disclosed. There are three major applications for the invention: roller pumps used for extracorporeal circulation, peristaltic pumps used for IV infusion, and peristaltic pumps for industry. The description of the invention hereinafter, makes extensive use of extracorporeal applications. This emphasis is made for description only. It should be understood that the described art is easily extended to the other aforementioned applications.
2. Description of the Prior Art.
The simplicity and availability of the standard roller pump have made it the choice for extracorporeal circulation. This pump is widely used in dialysis, routine cardiopulmonary bypass and long term pumping such as extracorporeal membrane oxygenators, (ECMO) and left and/or right heart bypass. The standard tubing used to pump blood with the roller pump is almost always polyvinyl chloride (PVC) with a nominal size from 1/4" ID with wall of 1/16" (Dialysis and pediatric cases) to 3/8" or 1/2" with a 3/32" wall. PVC tends to abrade easily, does not have the sufficient resilience at high pump speed to return to its natural diameter and tends to develop cracks along the edges that are bent repeatedly by the action of the pump. Manufacturers of roller pumps for extracorporeal circulation recommend that thick wall PVC tubing be used to overcome some of these problems. Use of thick walled tubing results in many disadvantages, among which are the following:
a) The roller pump maintains constant flow independent of clinically expected changes in inlet or outlet pressures. Thus, a decrease in blood supply at the pump inlet, without a concomitant decrease in pump speed, can cause excessive suction leading to air embolism, thrombosis and damage by the "venous" cannula to the patient's intima. Tests conducted by the inventor have shown that when starting at equilibrium using a polyvinyl chloride tube with 3/8" ID and 1/16" wall, a 30% drop in inlet flow caused the inlet pressure to drop from 0 to -250 mmHg; with a 3/32" wall tubing the pressure drop was from 0 to over -500 mmHg. The combination of constant flow and an arterial line that is accidentally clamped or kinked, or an arterial cannula that is positioned against the patient's intima, can generate excessive pressures at the outlet of the pump which at the extreme, can blow up a connector, tube, or an oxygenator.
To overcome these potential dangers, bubble oxygenators have blood level detectors which stop the pump when the blood level drops below a preset level, or a floating ball valve (SG-10 Ball Valve American Omni Medical Inc. Costa Mesa Calif. 92626), that closes when the blood level drops below the oxygenator outlet port. In closed systems such as ECMO or dialysis, easily collapsible bladders have been placed at the inlet to the pump such that at too high a suction, the bladder collapses actuating a microswitch which stops the pump. The pump restarts when the bladder refills. With current thick wall tubing this often results in very intermittent pumping because the slightest change in inlet flow causes the bladder to empty. Roller pumps with a microprocessor control servomotor may overcome some of these problems but their expense limits their use. When the roller pump is used during cardiac surgery for venting the left ventricle or for returning shed blood from the chest cavity it requires either the constant surveillance by a trained perfusionist to assure that excess suction does not occur, or the use of a suction relief valve such as the RLV-2100 " B" (American Omni Medical Inc., Costa Mesa Calif. 92626). This particular valve introduces air into the blood line which can increase hemolysis.
b) The power required to rotate the roller against the elastic force of the tubing wall is significant, requiring very powerful motors and indirect drive (e.g. gears) which, as will be described hereinafter, results in low pumping efficiencies.
c) The pumping rate is limited by the rotational speed of the motor and the resilience of the tubing returning the squeezed tubing wall to its natural open state. As will be described hereinafter, with the present invention of a thin wall tubing, a positive inlet pressure can better serve to maintain the tubing open.
d) At high rotational pump speed, the compression and decompression of the tubing wall generates heat and large flexture stresses that can change wall thickness resulting in a change in occlusiveness and limiting the pumping life of the tubing.
Studies by the inventor with various sized tubing with hardness of 60-75 Shore-A pumped with a roller pump indicate that normalized flow, expressed as a percentage of flow at atmospheric inlet pressure, was dependent on inlet pressure and independent of pump speed. It was also shown that the rate of change in normalized flow as a function of inlet pressure, referred to as flow sensitivity, (% flow)/mmHg, was directly related, and highly correlated (r.sup.2 &gt;98%) to the ratio of inside diameter (ID) to wall thickness (Wall) as follows: EQU Flow sensitivity=(1/3100)(ID/Wall).sup.3.24
or EQU ID/Wall=12(flow sensitivity).sup.0.309
Thus, a small change in ID/wall results in a large change in flow sensitivity.
U.S. Pat. Nos. 4,515,589 and 4,767,289 (manufactured by Sarns/3M Corp. of Ann Arbor Mich. as the "Safety Loop"), and 4,650,471, described devices to be used with the roller pump to regulate flow as a function of inlet pressure. Both of these devices utilize thin wall tubing (e.g. 0.375" ID with a 0.010" wall with ID/Wall=37.5) housed inside standard thick wall tubing. The thin wall tubing, which collapses easily in response to changes in inlet pressure, provides flow regulation and the thick wall housing provides mechanical support, prevents the thin wall tubing from getting tangled within the raceway of the roller pump and supports the thin walled tubing in the event that the outlet pressure causes the thin wall to herniate. The inventors disclosed the use of polyurethane as a possible material for the thin wall tubing in combination with a thick wall housing. Though excellent flow control can be achieved, the thick wall housing adds to the torque requirements of the pump.
The French Company Rhone-Poulenc sells the RP.01 Blood Pump that utilizes a silicone tube (10 mm ID by with a 1.3 mm wall, resulting in a relatively large ratio of ID to wall thickness of 7.7) that starts to collapse when the inlet pressure to the pump drops below -50 mmHg. As the inlet pressure drops from -50 to -100 mmHg the flow drops fivefold. The RP.01 pump also has a specially shaped roller that occludes only the central section of the tubing. This forms a channel at each edge of the tubing when the tubing is squeezed by the roller. Should air be aspirated by the pump, the low viscosity air would reduce the suction and stop blood flow. This pump tubing has four drawbacks: first, it can only be used with its own specially designed roller pump, second, there is no adjustment of the internal pressure about which the flow is controlled; third, the pump does not generate an outlet pressure much above 250 mmHg; and fourth the tubing used is silicone which easily spallates and has a short pumping life.
Two types of pump tubing made of polyurethane have been available commercially. First, a nominal 3/8" ID (measured 0.365") with a 0.040" wall pump chamber was manufactured by dip molding Biomer, a polyether-polyurethane (Ethicon, a Division of Johnson and Johnson, Sommerville N. J.). Additional mechanical strength was provided to this pump chamber with circumferential wrapping of fiberglass filament in a spiral manner about the wall of the Biomer. This Pump Chamber was limited in use because it was prohibitively expensive, over $100/ea in 1972. Second, a proprietary medical grade polyether-polyurethane (Pellethane 2363-80A) was extruded into standard size tubing (3/8" ID with 1/16" wall) for the roller pump (Tygothane by Norton Co., Akron Ohio 44309). This tubing proved to have longer pumping life and lower spallation (i.e. higher resistance to abrasion) than the standard tubing made of polyvinyl chloride (e.g. Tygon S-50-HL, Norton Co., Akron Ohio). However, Tygothane tubing was significantly harder then the standard polyvinyl chloride tubing (81 Shore A as compared to 65 Shore A) and caused the roller pumps to fail. It therefore never was used extensively and is no longer manufactured for roller pump use.
Theodore Kolobow of NIH reported [Transactions of American Society for Artificial Organs 15:172-177, 1969)] on a large bore thin walled tubing (15.3 mm ID, and 0.886.+-.0.25 mm wall, resulting in an ID/Wall=17.3) fabricated of Biomer by dip molding to give an " . . . exact clearance . . . " for setting the pump occulsion during long term pumping. This tubing was not used for flow control, which was achieved with a collapsible bladder in the venous line. Two years later Kolobow abandoned the thin wall tubing and substituted the aforementioned reinforced Biomer pump chamber made by Ethicon (ID/Wall=9). Currently Dr. Kolobow uses the aforementioned Tygothane tubing (1/2" ID by 1/16" wall) with a ID/Wall of 8.0. Tests by the inventor with that Tygothane tubing indicate that less than a 5% decrease in inlet flow results in a -400 mmHg decrease in inlet pressure.
Polyurethane was also used as an internal thin liner (about 0.010" thick) for standard polyvinyl chloride tubing (Bev-A-Line III). This tubing was coextruded and was claimed to have a much lower spallation and a negligible migration of plasticizers into blood. Its ID/Wall ratio equaled that of standard tubing and its major disadvantage was that the polyurethane liner separated from the polyvinyl chloride tubing when placed in the roller pump. This tubing is no longer manufactured for roller pump use.
Pierson et. al. [Transactions of American Society for Artificial Internal Organs 8:105-114, 1962] described a 1/4" to 1/2" ID thin walled tubing (0.020" wall, resulting in an ID/Wall as high as=18.7) fabricated of rubber latex to be used as an implantable or extracorporeal portable heart pump in conjunction with a high speed (RPM&gt;500) nonocclusive roller pump. Though this tubing was used for flow control as a function of inlet pressure, it was latex, which is manufactured by dipping, has a short pumping life, spallates, and could not support an outlet pressure higher than 100 mmHg. Such low pressure is incompatible with current extracorporeal technology. The motor used for the roller pump incorporated beveled gears which required more than 3 times the power required to pump blood at 2 l/min and an inlet pressure of 20 mmHg and outlet pressure of 100 mmHg. No reports were found of this system after the above mentioned reference. The inventor's experience with latex tubing suggests that such tubing may have very desirable flow characteristics as a function of inlet pressure but it cannot support clinically relevant outlet pressure [Tamari et.al. Transactions of American Society for Artificial Internal Organs 30:561-566, 1984]. Further, thin wall latex tubing cannot be used in a fully occlusive roller pump because the rollers would tend to stretch the elastic wall along the direction of rotation.
The poor choice of material may be why all the major manufacturers of roller pumps (e.g. Shiley, Cobe, 3M/Sarns) and of perfusion tubing (e.g. Baxter/Bentley, Norton, Shiley, Texas Medical Products) recommend using tubing #2 and #4 in Table 1 which have an ID/Wall ratio of 4.0. Cole-Parmer, a major supplier of peristaltic pumps and tubing for industry, does not have tubing with a ratio larger than 6 with most of the tubing having a ratio of 3 or less. The major reason for using a thick wall is to assure that pump flow does not change due to tubing fatigue resulting in a change from a round to oval shape cross section. This occurs because the tubing has insufficient resilience to withstand the repeated squeezing of the tube wall by the action of the pump. None of the manufacturers recommend using polyurethane tubing. Whenever commercially available thin wall tubing was used with the roller pump it was either mechanically supported (U.S. Pat. Nos. 4,515,589 and 4,767,289), reinforced with fiber filament (Ethicon) or a polyvinyl chloride liner (Bev-A-Line III), manufactured by dipping or could not support an outlet pressure greater than 100 mmHg.
The foregoing prior art may be summarized in table form as illustrated in Table 1 below. Tubing #4 and #5 are the most widely used tubing for roller pumps during adult cardiopulmonary bypass procedures.
TABLE 1 __________________________________________________________________________ Representative tubing used with peristaltic pumps. Tubing ID/ Hardness Strength Life Elong # Usage ID Wall Wall Material Shore A PSI Hours % __________________________________________________________________________ 1 CPB 3/16" 1/16" 3.0 PVC 60-70 2200 -- 350 2 Dial + CPB 1/4" 1/16" 4.0 PVC " -- " 3 CPB 3/8" 1/16" 6.0 PVC " 23 " 4 CPB 3/8" 3/32" 4.0 PVC " 55-123 " 5 CPB 3/8" 3/32" 4.0 PVC/PU " 55-123 " 6 CPB 1/2" 3/32" 5.3 PVC " -- " 7 RP.01 10 mm 1.3 mm 7.7 Silicone 1150 55 360 8 Ethicon* .365" .041" 9.0 Biomer 9 Kolobow** .500" .030" 17.3 Biomer 10 Pierson** 1/4-1/2" .020" 18.7 Latex 11 IV Pumps .100" .020" 5.0 PVC 60-75 12 Cole Parmer 1/16" 1/16" 1.0 PVC, Sil, 13 Cole Parmer 5/16" 1/16" 5.0 Neoprene 14 Invention .375" .030" 12.5 PEPU 72-80 4500 &gt;1700 650 __________________________________________________________________________ ID -- inside diameter, ID/wall -- ID divided by wall thickness, Hardness at room temperature, Strength -- tensile strength, Life -- pumping life obtained at 150-160 RPM and 500 mmHg outlet pressure at 30.degree. C., Elong % -- percent elongation at break. CPB -- Cardiopulmonary bypass, Dial -- Dialysis, PVC/PU Polyurethane lined PVC tubing, PEPU -- PolyetherPolyurethane. *Reinforced with spiral fiberglass filaments. **Made by dipping.
Independent of the tubing, damage to blood by the pumping action of the roller pump is mainly due to crushing of the cells between the opposite walls of the tubing. Blood damage is minimized by setting the pump nonocclusively; that is, allowing the space between the roller and the backplate of the pump head to be greater than twice the tubing wall thickness. The nonocclusive setting, typically allowing a 100 cm vertical column of blood to fall 2.5 cm/min, requires a very precise distance between the tubing rollers and the pump head raceway. Nonocclusive settings do not necessarily guarantee that the walls at the midsection of the tube do not contact; the cross section of squeezed tube within the roller head takes the shape of a figure "8" with the bent edges of the tubing last to occlude, and as will be hereinafter described in reference to FIG. 6(b). This is one of the main reasons that hemolysis rate for the roller pumps reported in the literature have been inconsistent. Another reason is that occlusion setting can also be inconsistent due to either nonuniform distance along the circumference of the housing backplate and the roller and variation in the wall thickness of standard perfusion tubing. The latter can initially be due to manufacturing (.+-.0.003" for one wall and.+-.0.006" for two walls) and, during pumping, due to temperature variations and material flow. The inability of the user to assure consistent low occlusion may be the reason that less occlusive setting causes less blood damage.
Some pump manufacturers (e.g. Cobe Laboratories) designed spring actuated rollers that push on the tubing, overcoming the elasticity of the tubing wall, and provide uniform occlusion independent of the aforementioned distance and wall variations. The disadvantages of this system are that the tubing is set occlusively, increasing blood damage, and that it is difficult to use with the more rigid and thicker tubing currently used for the heart-lung machine.
Currently, motors for roller pumps generate a peak torque of 70 in-lb which is equivalent to an outlet pressure of over 11,000 mmHg, or pressure that is over 35 times greater than the maximum pressure used clinically. This excessive torque can deform 3/8" ID with 3/32" wall PVC tubing placed in the roller pump if the outlet to that tubing is clamped.
It would be of great advantage to have tubing for peristaltic pumps that can provide one or more of the following characteristics: long pumping life, low spallation, self regulation, low power requirement, assure easy and accurate setting of occlusion, support an outlet pressure compatible with its use without affecting pumping, and reduce hemolysis. In addition it would also would be advantageous to have the tube made of material that is extrudable, allow inexpensive means to measure the pumped fluid pressures at the inlet and outlet of the pump without direct contact with pumped fluid, and allow flow control within the limits set for the inlet and outlet pressure of the pump. Such characteristics can be achieved with the present invention of extruded polyurethane tubing with at least one thin wall section when used with a conventional roller pump.
As part of an overall system to improve peristaltic pumping of blood, the present invention features a new innovative roller pump that takes advantage of the thin wall tubing as the pump tubing. As will be described hereinafter, the pump has a significantly higher pumping efficiency, lower power requirements, inlet and outlet pressure monitoring capabilities, flow regulation. The pumps would be safer, compact, require smaller batteries, lower power and be computerized and user friendly.
The thin wall tubing with the aforementioned advantages can also be used with peristaltic pumps used for intravenous fluid administration such as Flo-Gard 2000 made by Baxter Health Care or as described by many U.S. Patents, as for example, U.S. Pat. No. 4,702,675.
Presently there are two types of infusion pumps, the peristaltic pump using IV tubing, and the piston type pump that requires an expensive cassette (e.g. Flo-Gard 8500 made by Baxter-Travenol). The peristaltic pump is easy to control and use; however, the tubing used with it is almost exclusively made of polyvinyl chloride (PVC) with some limited use of silicone. This has the following disadvantages:
1. Both tend to spallate, abrading material from the tubing wall thereby releasing particles into the pumped fluid.
2. PVC loses its resilience over time (fatigue), thereby not returning to its original shape and resulting in an undesirable decrease in flow.
3. Relatively thick wall tubing is used (e.g. 0.100" ID with 0.140" OD resulting in a ID/Wall=5). Thick wall tubing requires peristaltic pumps that have higher torque motors to overcome the higher resistance to wall compression with most of the pump's energy being wasted on compressing the tubing wall. The relatively thick wall prevents accurate monitoring of the IV infusion pressure, which is used to alarm upon occlusion of said IV line by comparing that pressure to the pressure required for infusion. The thick wall also prevents accurate measurements of inlet pressure which is used to provide a low filling pressure alarm to indicate lower flow.
These problems were solved with the use of a cassette-type positive displacement volumetric pump (e.g. Baxter-Travenol's Flo-Gard 8500) but at a considerable expense: the cassette increases the costs of IV sets by 3 to 4 times. It would be of great clinical advantage to be able to overcome the above problems and measure the infusion pressure accurately with the less expensive disposable tubing used with peristaltic pumps such as Flo-Gard 6200.
Another positive displacement infusion system for IV administration which attempts to reduce power consumption of the pump and reduce costs is described by U.S. Pat. No. 4,846,637. This invention suggests using a specially formed pumping chamber comprised of an elastomeric material such as urethane or silicone with a first nondeformable portion of substantial thickness that provides strength and support during compression of said chamber and a second relatively thin deformable portion which forms an elliptical fluid conduit. Low power requirements are attributed to the thin wall and to its elliptical shape.